Method of reducing spatial extent of gradient coil current feeding connectors

ABSTRACT

A magnetic resonance (MR) imaging system includes: a main magnet configured to generate a volume of magnet field suitable for forming MR imaging; a transmit radio frequency (RF) coil assembly configured to transmit at least one RF signal; a receive radio frequency (RF) coil assembly configured to, in response to the at least one RF pulse, receive MR signals; a gradient coil assembly comprising (i) windings of coils arranged in a radial layer, and (ii) a first set of connectors embedded in the radial layer to reduce a radial extent occupied by the gradient coil assembly, the first set of electrical connectors configured to drive the windings of coils; and a control unit coupled to the transmit RF coil assembly, the receive RF coil assembly, and the gradient coils, the control unit configured to synchronously operate the gradient coil assembly, the transmit coil assembly, and the receive coil assembly.

BACKGROUND

The present disclosure relates to magnetic resonance imaging.

SUMMARY

In one aspect, some implementations provide a magnetic resonance (MR)imaging system, including: a housing having a bore shaped and sized toaccommodate at least a portion of a subject; a main magnet accommodatedby the housing and configured to generate a volume of magnet fieldsuitable for forming MR imaging over a region located within the boreand covered by the volume of magnetic field; a transmit radio frequency(RF) coil assembly configured to transmit at least one RF signal intothe portion of the subject; a receive radio frequency coil assemblyconfigured to, in response to the at least one RF pulse, receive MRsignals emitted from the portion of the subject; a gradient coilassembly comprising (i) windings of coils arranged in a radial layer,and (ii) a first set of connectors embedded solely in the radial layerto reduce a radial extent occupied by the gradient coil assembly, thefirst set of electrical connectors configured to drive the windings ofcoils with sufficient currents suitable to generate perturbations to thevolume of magnet field such that the MR signals encode an MR image overthe region and in accordance with the generated perturbations; and acontrol unit coupled to the transmit RF coil assembly, the receive RFcoil assembly, and the gradient coils, the control unit configured tosynchronously operate the gradient coil assembly, the transmit coilassembly, and the receive coil assembly.

Implementations may include one of more of the following features.

The radial layer may include a depressed area where the first set ofelectrical connectors are inlaid such that the first set of electricalconnectors are, at least in part, radially constrained between theradial layer's outer surface and inner surface. The first set ofelectrical connectors may be separated from the underlying radial layerby electrical insulation material. The gradient coil assembly may beconfigured to generate perturbations to the volume of magnet field alonga longitudinal axis of the bore or transverse to the longitudinal axisof the bore.

The MR imaging system may further include shield gradient coils arrangedradially outside the windings of coils of the gradient coil assembly andconfigured to generate a varying magnetic field that attenuates adifferent varying magnetic field generated by the gradient coil assemblyoutside of the bore.

The MRI system may further include shimming coils located on the radiallayer or a different radial layer, wherein the shimming coils areconfigured to improve a homogeneity of the magnetic field within theregion for MR imaging, wherein the shimming coils are driven by shimmingcurrents provided through a second set of electrical connectors locatedon the radial layer or the different radial layer, wherein the secondset of electrical connectors are different from the first set ofelectrical connectors.

The MR imaging system may further include cooling structures inlaid inthe radial layer and configured to provide cooling when the windings ofcoils of the gradient coil assembly are driven with sufficient currentssuitable to generate perturbations to the volume of magnet field.

The gradient coil assembly may be coupled to a gradient amplifierthrough the first set of electrical connectors. The first set ofelectrical connectors may be configured to receive currents up to rootmean square (rms) 400 A continuous.

The gradient amplifier may drive the gradient coil assembly through thefirst set of electrical connectors with sufficient currents suitable togenerate perturbations to the volume of magnet field that are along andtransverse a longitudinal axis of the bore. The control unit may befurther configured to reconstruct a magnetic resonance (MR) image basedon the MR signals.

The MR imaging system may further include a display on which the MRimage is presented.

In another aspect, implementations may include method for manufacturingan MRI system, the method including: configuring a housing that includesa bore shaped and sized to accommodate at least a portion of a subject;arranging a main magnet to be accommodated by the housing andconfiguring the main magnet to generate a volume of magnet fieldsuitable for forming MR imaging over a region located within the boreand covered by the volume of magnetic field; configuring a transmitradio frequency (RF) coil assembly capable of transmitting at least oneRF signal into the portion of the subject; configuring a receive radiofrequency (RF) coil assembly capable of, in response to the at least oneRF pulse, receiving MR signals emitted from the portion of the subject;configuring a gradient coil assembly by: arranging windings of coilsarranged in a radial layer, and embedding a first set of electricalconnectors solely in the radial layer to reduce a radial extent occupiedby the gradient coil assembly and configuring the first set ofelectrical connectors to drive the windings of coils with sufficientcurrents suitable to generate perturbations to the volume of magnetfield such that the MR signals encode an MR image over the region and inaccordance with the generated perturbations; and configuring a controlunit to be coupled to the transmit RF coil assembly, the receive RF coilassembly, and the gradient coils such that the control unit is capableof synchronously operating the gradient coil assembly, the transmit coilassembly, and the receive coil assembly.

Implementations may include one or more the following features.

Embedding the first set of electrical connectors comprises inlaying thefirst set of electrical connectors in a depressed area on the radiallayer such that the first set of electrical connectors are, at least inpart, radially constrained between the radial layer's outer surface andinner surface.

Embedding the first set of electrical connectors in a depressed area maycomprise soldering the first set of electrical connectors to provide anelectrical connection to the windings of coils, and wherein theelectrical connection is formed in the depressed area such that thefirst set of electrical connectors do not extend above the outer surfaceof the radial layer.

The method may further include: arranging shield gradient coils radiallyoutside the windings of coils of the gradient coil assembly; andconfiguring the shield gradient coil to be capable of generating avarying magnetic field that attenuates a different varying magneticfield generated by the gradient coil assembly outside of the bore.

The method may further include: arranging shimming coils to bepositioned on the inside or the outside of the radial layer; andconfiguring the shimming coils to improve a homogeneity of the magneticfield within the region for MR imaging when the shimming coils aredriven by shimming currents provided through a second set of electricalconnectors different from the first set of electrical connectors.

The method may further include: arranging cooling structures to bepositioned on the radial layer or on a different radial layer underneaththe radial layer; and configuring the cooling structures such thatcooling is effectuated when the windings of coils of the gradient coilassembly are driven with sufficient currents suitable to generateperturbations to the volume of magnet field.

The method may further include: coupling the gradient coil assembly to agradient amplifier through the first set of electrical connectorscapable of receiving currents up to root mean square (rms) 400 Acontinuous.

The method may further include: coupling the control unit to a displaysuch that when the control unit has reconstructed an magnetic resonance(MR) image based on the MR signals, the display is configured to presentthe reconstructed MR image.

The details of one or more aspects of the subject matter described inthis specification are set forth in the accompanying drawings and thedescription below. Other features, aspects, and advantages of thesubject matter will become apparent from the description, the drawings,and the claims.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1A shows a perspective view of an example of a magnetic resonanceimaging (MRI) system with a solenoid magnet.

FIG. 1B shows a cross-sectional illustration of the example of amagnetic resonance imaging (MRI) system.

FIGS. 2A to 2B illustrates examples of reducing the spatial extent of agradient coil assembly.

FIGS. 3A to 3C illustrates perspective views of reducing the radialfootprint of current feeding connectors on a prototype gradient coilassembly.

FIGS. 4A to 4C illustrate examples of reducing the spatial extent ofcooling structures in a magnetic resonance imaging (MRI) system.

Like reference symbols in the various drawings indicate like elements.

DETAILED DESCRIPTION

Various embodiments and aspects of the disclosure will be described withreference to details discussed below. The following description anddrawings are illustrative of the disclosure and are not to be construedas limiting the disclosure. Numerous specific details are described toprovide a thorough understanding of various embodiments of the presentdisclosure. However, in certain instances, well-known or conventionaldetails are not described in order to provide a concise discussion ofembodiments of the present disclosure.

Gradient coils are employed by magnetic resonance imaging (MRI) systems.An example of a gradient coil assembly on an MRI system may provideperturbation of the main magnetic field during an MRI imaging sequenceso that MR signals may encode proton resonance emitted from variousparts of a subject during an MR scan and an MR image may bereconstructed based on the MR signals. Gradient coils tend to bephysically large and massive components of the MRI system with manylayers. Yet, gradient coil designs typically do not have abundant layerspace available for the internal structures. For a cylindrical gradientcoil, the outer diameter should be as small as possible to accommodate,for example, a smaller main magnet for cost and siting advantages whilethe inner diameter should be as large as possible to accommodate, inparticular, a larger patient bore for comfort and access. When balancingsuch competing design goals, the performance of the gradient coil can beimproved by having the primary innermost gradient coil layers located asclose as possible to the inner patient-side surface of the gradient coiland the shielding gradient coil layers be as close as possible to theouter magnet-side surface of the gradient coil. For a gradient coillayer to operate, a method of electrical current supply and returnshould be established to have at least one of the feeding electricalconnectors overlap with the coil layer at some distance from its outeredge. If such an electrical connector takes up layer space, then theavailable space for placing other components is reduced, furthercompounding the space limitations. This situation can be problematicespecially where the adjacent layer has features which have aperpendicular direction to the electrical connector, for example, aspiral of cooling hoses. The same issues generally plague the design anddevelopment of electromagnet systems with multiple layers. In otherwords, such issues may not be limited to gradient coils on an MRIsystem.

Implementations of the disclosure endeavor to reduce or even eliminatelayer space used by feeding electrical connectors by providing a channelor depression in, for example, the gradient coil layer in which thefeeding electrical connectors are embedded such that its height abovethe gradient coil layer is reduced or substantially eliminated, asexplained below.

FIGS. 1A-1B show a perspective view and a cross-sectional view of anexample of a magnetic resonance imaging (MRI) system 100 in which asolenoid magnet 105 is provided in a cylindrical shape with an innerbore 101. Coil assembly 107, including transmit RF coil 106 and gradientcoil 104, is provided within solenoid magnet 105. Coil assembly 107 maygenerally be shaped as an annular structure and housed within the innerbore of solenoid magnet 105. In some implementations, coil assembly 107is annular and only includes gradient coil 104. In theseimplementations, annular coil assembly does not include transmit RF coil106 or any receiver coil. For these implementations, radio frequency(RF) excitation pulses are, for example, transmitted by local coils forimaging the head region 102 of patient 103. In one instance, a head coilin a birdcage configuration is used for both transmitting RF excitationpulses and receiving MR signals for imaging the subject. An example of abirdcage configuration is shown in FIG. 2, which may be used as atransmit-only RF coil assembly. In this case, a birdcage coil may beused for transmitting an RF excitation pulse into the subject and aparallel array coil configuration is used for receiving MR signals inresponse.

In some implementations, shimming coils 109 are housed within thegradient coil assembly. Shimming coils 109 are powered by a group ofpower amplifiers. In some cases, the power amplifiers are housed in acontrol room and are connected to shimming coils 109 to provide shimmingof the magnetic field within inner bore 101. In driving shimming coils109, power amplifiers may be controlled by a control unit that generallyincludes one or more processors as well as programming logic toconfigure the power amplifiers. In some instances, the control unit ishoused in a control room separate from the solenoid magnet 105 of theMRI system 100. The control unit is further configured to reconstruct amagnetic resonance (MR) image based on the MR signals. The MRI systemmay include a display on which the reconstructed MR image is presented,for example, to a human operator. The driving current for shimming coils109 may be in the range of hundreds of miliamperes and generally do notexceed 20 ampere. Further, shimming coils 109 may not require activecooling using circulating coolant. In these implementations, an array ofshimming coils can be used to provide adjustment to the field strengthwithin the inner bore 101 such that the magnet field within the innerbore 101 becomes more homogenous.

The implementations described in this disclosure may be adapted forintraoperative MRI, and MRI systems for use in an emergency roomsetting. Such MRI systems may include a smaller and more compact boresize magnet compared to the magnets from conventional whole bodyscanners. One consequence of a smaller bore magnet is that, the volumeof uniform magnetic field suitable for imaging (e.g., with fieldinhomogeneity below a defined threshold) may not cover all areas ofinterest.

Referring to FIGS. 1A to 1B, the transmit RF coil 106 is a resonantstructure that excites the sample during a magnetic resonance imaging(MRI) acquisition. The resonant frequency of the transmit RF coil 106 istuned to the Larmor frequency for the nuclei of interest and fieldstrength of the MRI system 100. The input impedance of the tuned RF coilis then transformed at the coil input to match the amplifiercharacteristic impedance (typically 50 Ω).

Gradient coil 104 generally provides field gradients in more than onedirection, such as, for example, all three orthogonal spatialdirections. Thus, gradient coil 104 may refer to three sets of coils,each configured to generate field fluctuations in a respective directionfor the main field in the inner bore of the solenoid magnet 105. Suchfield fluctuations may cause magnetizations from various spatiallocations to experience precessions at different frequencies, enablingencoding of spatial information of the magnetizations through RFexcitation pulses.

In more detail, gradient coils can be physically large and massivecomponents of the MRI system with many layers. A gradient coil'sinternal structure can include space for electromagnets that make up theX, Y and Z gradient coil axes, higher-order shim coil axes, passive shimtrays, structures for mechanical support, and tubing or components ofthe cooling system. Gradient coil design can be limited by availablelayer space for these structures. For a cylindrical gradient coil, theouter diameter should be made as small as possible to accommodate asmaller main magnet for cost and siting advantage while the innerdiameter should be kept as large as possible to accommodate a largerpatient bore for comfort and access. These preferences are analogous forother gradient coil geometries, for example planar gradient coils. Suchcompeting design goals must be balanced by the performance of thegradient coil which is improved by having the primary innermost gradientcoil layers be as close as possible to the inner patient-side surface ofthe gradient coil and the shielding gradient coil layers be as close aspossible to the outer magnet-side surface of the gradient coil. If theindividual gradient coil layers could be made in such a way as to reducethe radial space while preserving the physical extents of the entirestructure, then shield gradient coil layers could be placed farther fromprimary gradient coil layers, thus increasing possible performance, ormore space would be available in the internal structure of the gradientcoil to accommodate, for example, additional shim coil layers or coolinglayers.

A method of electrical current supply and return may be used to operatethe gradient coil 104. This method may utilize at least one of thefeeding electrical connectors for providing the current supply. On aconventional gradient coil system, this feeding electrical connectoroverlaps with the gradient coil layer at least some distance from theouter edge of gradient coil layer. Because such an electrical connectortakes up layer space on the conventional gradient coil system, space forplacing other components is thus further reduced. This situation isproblematic especially where the adjacent layers possess featuresarranged in a direction substantially perpendicular to the electricalconnector, for example a spiral of cooling hose. The same issuesgenerally plague the design and development of electromagnet withmultiple layers, not just gradient coils for MRI systems.

To use less layer space, gradient coils may be designed to have allelectrical contacts for each loop occur at the end of the coil wherethere may be more space to interleave connectors, but the drawback withthis approach is that it imposes a design restriction which can limitperformance of the gradient coil system in terms of the range and extentof perturbations to the main magnetic field that can be realized. Asexplained, the perturbations to the main magnetic field determine theencoding of nuclei magnetic resonance signals emitted from a subjectplaced inside the magnetic field to form an MR image.

Another approach is to orient layer stacking such that the electricalconnectors lie in between gaps of adjacent layer structures. Thisapproach can be equally undesirable because it is not always feasiblewithout imposing design restrictions.

In yet another approach, the electrical connectors can be brought backin other layer spaces where there may be room farther away from therespective layer, but the drawback is that this approach reduces theelectromagnetic cancellation of the magnetic field deviation introducedby the jog of the wire pattern of the gradient coil layer.

In sum, most conventional systems incorporate gradient coils that aredesigned with connectors occupying radial space, thus increasing theextent of the physical space that the gradient coil layers wouldotherwise occupy.

Referring to FIGS. 2A-2B, examples of embodiments that embed or flattenthe current feeding connectors in a gradient coil assembly in a mannerthat achieves radial space reduction and preserves gradient systemperformance. The perspective views in the top panels demonstrate generalexamples of spiral-shape electromagnets (‘thumbprints’) viewed from thetop down or bottom up. These perspectives may be flat (as shown in thefigures) or they may be curved, similar to the view shown in FIGS. 3A to3C. These features are generally curved for cylindrical-geometry MRIgradients. For other geometries of MRI systems, such as permanent magnetdesigns with planar gradients, these kinds of electromagnets aredesigned to be kept flat.

Radial layer 206 represents an outer surface on a cylindrical structurethat houses windings of coils, as illustrated by the spiral lines.Radial layer 206 may be made from, for example, carbon fiber/epoxycomposite material. These coil windings form the backbone of a gradientcoil assembly from which perturbations to the main magnetic field aregenerated for the purposes of encoding a nucleic magnetic resonancesignal. Here, the cylindrical structure has a first end 206E1 and asecond end 206E2. A first current feeding connector 202 is located atthe first end 206E1 while a second current feeding connector 204 islocated at the second end 206E2. Currents can flow into the windings ofcoils through one of the current feeding connectors and then flow out ofthe windings of coils through the other of the current feedingconnectors.

Briefly referring to FIGS. 3A to 3C, a prototype gradient coil assembly300 is shown. In one illustration from FIG. 3A, earlier examples mayplace current feeding connectors 302 and 304 on the radial layer 306.While these examples may be expeditiously configured, in other examples,as shown in FIGS. 3B to 3C, a first current feeding connector 302 isembedded in radial layer 306 at a first end 306E1 while a second currentfeeding connector 304 is embedded in radial layer 306 at a second end306E2. A gradient coil, for example, an x-axis gradient coil or a y-axisgradient coil may be, placed at region 308 of radial layer 306. Currentsmay flow into the gradient coil in region 308 through the first currentfeeding connector 302 and then flow out of the gradient coil through thesecond current feeding connector 304. Comparing and contrasting FIGS. 3Ato 3C reveals the savings in the radial extent as occupied by thecurrent feeding connector.

Returning to FIG. 2A, the bottom panel shows a cross-sectional view toillustrate the embedded placement of the current feeding connectors 202and 204. Here, the current feeding connectors 202 and 204 haverectangular profiles and are embedded into the radial layer 206 in amanner that neither of the current feeding connectors extend above theouter surface of radial layer 206.

FIG. 2B shows a similar radial layer 216 on an outer surface on acylindrical structure that houses the windings of coils, as illustratedby the spiral lines. Again, these coil windings form the backbone of agradient coil assembly from which perturbations to the main magneticfield are generated, as illustrated, for example, in the perspectiveview of FIG. 3. Here, the cylindrical structure has a first end 216E1and a second end 216E2. A first current feeding connector 212 is locatedat the first end 216E1 while a second current feeding connector 204 islocated at the second end 216E2. Currents can flow into the windings ofcoils through one of the current feeding connectors and then flow out ofthe windings of coils through the other of the current feedingconnectors. The bottom panel shows a cross-sectional view to illustratethe embedded placement. A side-by side comparison with FIG. 2A revealsthat, the current feeding connectors 212 and 214 have rectangularprofiles and may be inlaid on the outer surface of the radial layer 216by virtue of a reduced height but increased lateral span. Moreparticularly, the cross cuts of FIG. 2B break up large area surfaces toreduce the formation of large loop eddy currents, since these connectorsoperate in a rapidly changing magnetic field environment by virtue ofbeing inside a gradient coil. Without the cross cuts, these connectorswould experience greater eddy current heating. The cross cuts also helpconstrain the current flow to be in the center of the connector, asdesigned. The arrangement in FIG. 2B maintains a high cross section areafor most of the connector, compared when the connector is narrowed, thushaving substantially similar Ohmic heating characteristics to a thickerand narrower connector. More details on the particular design andrationale can be found in a co-pending PCT Patent Application,WO2017077368 A1, which is incorporated by reference in its entirety.

In these illustrations, the gradient coil may be coupled to a gradientamplifier through the current feeding connectors. Indeed, these currentfeeding connectors are capable of handling currents up to root meansquare (RMS) 400 A in continuous mode. The gradient amplifier drives thegradient coil assembly through the current feeding connectors withsufficient currents suitable to generate perturbations to the volume ofmagnet field that are along and transverse a longitudinal axis of thebore.

Referring to FIGS. 4A to 4C, an illustration shows another similarradial layer 426 that houses cooling structures such as cooling lines428 and 429 in the same radial layer as the gradient coil currentfeeding connectors 422 and 424. The cooling lines 428 and 429 can bearranged in proximity to the heat-generating radial layer 426 housing,for example, a gradient coil assembly. The gradient coil assembly mayinclude current feeding connectors 422 and 424 that are respectivelylocated at first end 426E1 and second end 426E2 with respect to radiallayer 426. The proximity in placement allows heat to be extracted fromradial layer 426 through conduction and convection. The cooling lines428 and 429 may be embedded or inlayed in channels cut into radial layer426. Coolant (e.g., water) may flow into inlet 428C1 and out of outlet428C2. Likewise, 429C1 is a coolant inlet and 429C2 is a coolant outlet.The coolant may flow from other tubing or manifolds placed inside thegradient coil assembly in radial layer 426 or outside the gradient coilassembly in radial layer 426. While in this illustration the coolinglines (428 and 429) and the gradient coil current feeding connectors(422 and 424) are placed in the same radial layer 426, in someconfigurations, the cooling lines and the gradient coil current feedingconnectors 426 may be placed in different but adjacent radial layers,each of which configured with reduced radial profile to accommodateefficient spatial stacking.

The illustration also shows one tubing represented by two tubingstructures representing cooling lines, namely 428 and 429, to providecooling during operation of the magnetic resonance imaging system. Thefirst loop is from a first end 428C1 to a second end 428C2. The secondloop is from a third end 429C1 to a fourth end 429C2. The bottom panelat FIG. 4A shows a cross sectional view of the radial layer 426 in oneembodiment where tubing structures (representing cooling lines 428 and429) have circular profiles and are embedded in radial layer 426. Insome cases, the tubing structures 428 and 429 may be made of heatconducting material (with some electrical insulation provided betweenthe cooling tube structures and the radial layer 426). In other cases,the tubing may be made of non-conducting material. In one example, asillustrated by FIG. 4A, channels may be milled partially into the radiallayer 426 to inlay these tubes fully or partially within the radiallayer 426. In another example, the radial layer 426 may include a layerof electrically insulating material bonded to the conductor (i.e. byepoxy) and channels for the coolant tubing may be milled or formed intothis insulating material, or into both the insulating material andpartially into the conductor of the electromagnet.

The tubes may also be formed into a non-circular cross section (i.e. bymeans of a press to make oval shaped cross section, or by extruding thetubing material in a non-circular shape) so as to reduce further theradial extent of the cooling tubes without substantially reducing thecoolant flow cross section. The cooling tubes would need to be made of amalleable material that remains in the new cross sectional shape (i.e.copper or some kinds of plastic). The cooling structures may be placedon the inside or outside of the curved face of, for example, radiallayer 426 to provide cooling when, for example, the gradient coils aredriven with sufficient currents for MRI imaging. Here, inlaying thecooling tubes are inspired by the same motivations as inlaying theelectrical connecting pieces because the advantages derived arecomparable. Indeed, the inventive concept of embedding current feedingconnectors may be extended to spur development of shimming coils as wellas cooling structures based on similar motivations and advantages. Inthe case of shimming coils, the shimming coils do not need to be on theparticular radial layer. In fact, the shimming coils can be located ondifferent radial layers.

As used herein, the terms “comprises” and “comprising” are to beconstrued as being inclusive and open ended, and not exclusive.Specifically, when used in the specification and claims, the terms“comprises” and “comprising” and variations thereof mean the specifiedfeatures, steps or components are included. These terms are not to beinterpreted to exclude the presence of other features, steps orcomponents.

As used herein, the term “exemplary” means “serving as an example,instance, or illustration,” and should not be construed as preferred oradvantageous over other configurations disclosed herein.

As used herein, the terms “about” and “approximately” are meant to covervariations that may exist in the upper and lower limits of the ranges ofvalues, such as variations in properties, parameters, and dimensions. Inone non-limiting example, the terms “about” and “approximately” meanplus or minus 10 percent or less.

The specific embodiments described above have been shown by way ofexample, and it should be understood that these embodiments may besusceptible to various modifications and alternative forms. It should befurther understood that the claims are not intended to be limited to theparticular forms disclosed, but rather to cover all modifications,equivalents, and alternatives falling within the spirit and scope ofthis disclosure.

What is claimed is:
 1. A magnetic resonance (MR) imaging system,comprising: a housing having a bore shaped and sized to accommodate atleast a portion of a subject; a main magnet accommodated by the housingand configured to generate a volume of magnet field suitable for formingMR imaging over a region located within the bore and covered by thevolume of magnetic field; a transmit radio frequency (RF) coil assemblyconfigured to transmit at least one RF signal into the portion of thesubject; a receive radio frequency (RF) coil assembly configured to, inresponse to the at least one RF pulse, receive MR signals emitted fromthe portion of the subject; a gradient coil assembly comprising (i)windings of coils arranged in a radial layer, and (ii) a first set ofconnectors embedded solely in the radial layer to reduce a radial extentoccupied by the gradient coil assembly, the first set of electricalconnectors configured to drive the windings of coils with sufficientcurrents suitable to generate perturbations to the volume of magnetfield such that the MR signals encode an MR image over the region and inaccordance with the generated perturbations; and a control unit coupledto the transmit RF coil assembly, the receive RF coil assembly, and thegradient coils, the control unit configured to synchronously operate thegradient coil assembly, the transmit coil assembly, and the receive coilassembly.
 2. The MRI system of claim 1, wherein the radial layercomprises a depressed area where the first set of electrical connectorsare inlaid such that the first set of electrical connectors are, atleast in part, radially constrained between the radial layer's outersurface and inner surface.
 3. The MRI system of claim 1, wherein thefirst set of electrical connectors are separated from the underlyingradial layer by electrical insulation material.
 4. The MRI system ofclaim 1, wherein the gradient coil assembly is configured to generateperturbations to the volume of magnet field along a longitudinal axis ofthe bore or transverse to the longitudinal axis of the bore.
 5. The MRIsystem of claim 1, further comprising: shield gradient coils arrangedradially outside the windings of coils of the gradient coil assembly andconfigured to generate a varying magnetic field that attenuates adifferent varying magnetic field generated by the gradient coil assemblyoutside of the bore.
 6. The MRI system of claim 1, further comprising:shimming coils located on the radial layer or a different radial layer,wherein the shimming coils are configured to improve a homogeneity ofthe magnetic field within the region for MR imaging, wherein theshimming coils are driven by shimming currents provided through a secondset of electrical connectors located on the radial layer or thedifferent radial layer, wherein the second set of electrical connectorsare different from the first set of electrical connectors.
 7. The MRIsystem of claim 1, further comprising: cooling structures inlaid in theradial layer and configured to provide cooling when the windings ofcoils of the gradient coil assembly are driven with sufficient currentssuitable to generate perturbations to the volume of magnet field.
 8. TheMRI system of claim 1, wherein the gradient coil assembly is coupled toa gradient amplifier through the first set of electrical connectors. 9.The MRI system of claim 8, wherein the first set of electricalconnectors are configured to receive currents up to root mean square(rms) 400 A continuous.
 10. The MRI system of claim 1, wherein thegradient amplifier drives the gradient coil assembly through the firstset of electrical connectors with sufficient currents suitable togenerate perturbations to the volume of magnet field that are along andtransverse a longitudinal axis of the bore.
 11. The MRI system of claim1, wherein the control unit is further configured to reconstruct amagnetic resonance (MR) image based on the MR signals.
 12. The MRIsystem of claim 1, further comprising: a display on which the MR imageis presented.
 13. A method for manufacturing an MRI system, the methodcomprising: configuring a housing that includes a bore shaped and sizedto accommodate at least a portion of a subject; arranging a main magnetto be accommodated by the housing and configuring the main magnet togenerate a volume of magnet field suitable for forming MR imaging over aregion located within the bore and covered by the volume of magneticfield; configuring a transmit radio frequency (RF) coil assembly capableof transmitting at least one RF signal into the portion of the subject;configuring a receive radio frequency (RF) coil assembly capable of, inresponse to the at least one RF pulse, receiving MR signals emitted fromthe portion of the subject; configuring a gradient coil assembly by:arranging windings of coils arranged in a radial layer, and embedding afirst set of electrical connectors solely in the radial layer to reducea radial extent occupied by the gradient coil assembly and configuringthe first set of electrical connectors to drive the windings of coilswith sufficient currents suitable to generate perturbations to thevolume of magnet field such that the MR signals encode an MR image overthe region and in accordance with the generated perturbations; andconfiguring a control unit to be coupled to the transmit RF coilassembly, the receive RF coil assembly, and the gradient coils such thatthe control unit is capable of synchronously operating the gradient coilassembly, the transmit coil assembly, and the receive coil assembly. 14.The method of claim 13, wherein embedding the first set of electricalconnectors comprises inlaying the first set of electrical connectors ina depressed area on the radial layer such that the first set ofelectrical connectors are, at least in part, radially constrainedbetween the radial layer's outer surface and inner surface.
 15. Themethod of claim 14, wherein embedding the first set of electricalconnectors in a depressed area comprises soldering the first set ofelectrical connectors to provide an electrical connection to thewindings of coils, and wherein the electrical connection is formed inthe depressed area such that the first set of electrical connectors donot extend above the outer surface of the radial layer.
 16. The methodof claim 13, further comprising: arranging shield gradient coilsradially outside the windings of coils of the gradient coil assembly;and configuring the shield gradient coil to be capable of generating avarying magnetic field that attenuates a different varying magneticfield generated by the gradient coil assembly outside of the bore. 17.The method of claim 13, further comprising: arranging shimming coils tobe positioned on the inside or the outside of the radial layer; andconfiguring the shimming coils to improve a homogeneity of the magneticfield within the region for MR imaging when the shimming coils aredriven by shimming currents provided through a second set of electricalconnectors different from the first set of electrical connectors. 18.The method of claim 13, further comprising: arranging cooling structuresto be positioned on the radial layer or on a different radial layerunderneath the radial layer; and configuring the cooling structures suchthat cooling is effectuated when the windings of coils of the gradientcoil assembly are driven with sufficient currents suitable to generateperturbations to the volume of magnet field.
 19. The method of claim 13,further comprising: coupling the gradient coil assembly to a gradientamplifier through the first set of electrical connectors capable ofreceiving currents up to root mean square (rms) 400 A continuous. 20.The method of claim 13, further comprising: coupling the control unit toa display such that when the control unit has reconstructed an magneticresonance (MR) image based on the MR signals, the display is configuredto present the reconstructed MR image.